A biosensor (or biological sensor) is an analytical device including a biological element and a transducer that converts a biological response into an electrical signal. Certain biosensors involve a selective biochemical reaction between a specific binding material (e.g., an antibody, a receptor, a ligand, etc.) and a target species (e.g., molecule, protein, DNA, virus, bacteria, etc.), and the product of this highly specific reaction is converted into a measurable quantity by a transducer. Other sensors may utilize a non-specific binding material capable of binding multiple types or classes of molecules or other moieties that may be present in a sample. The term “functionalization material” may be used herein to generally relate to both specific and non-specific binding materials. Transduction methods used with biosensors may be based on various principles, such as electrochemical, optical, electrical, acoustic, etc. Among these, acoustic transduction offers a number of potential advantages, such as being real time, label-free, and low cost, as well as exhibiting high sensitivity.
An acoustic wave device employs an acoustic wave that propagates through or on the surface of a specific binding material, whereby any changes to the characteristics of the propagation path affect the velocity and/or amplitude of the wave. Acoustic wave devices are commonly fabricated by micro-electromechanical systems (MEMS) fabrication techniques, owing to the need to provide microscale features suitable for facilitating high-frequency operation. Presence of functionalization material on or over an active region of an acoustic wave device permits an analyte to be bound to the functionalization material, thereby altering the mass being vibrated by the acoustic wave and altering the wave propagation characteristics (e.g., velocity, thereby altering resonance frequency). Changes in velocity can be monitored by measuring the frequency, magnitude, and/or phase characteristics of the acoustic wave device and can be correlated to a physical quantity being measured.
In the case of a piezoelectric crystal resonator, an acoustic wave may embody a bulk acoustic wave (BAW) propagating through the interior (or “bulk”) of a piezoelectric material. BAW devices typically involve transduction of an acoustic wave using electrodes arranged on opposing top and bottom surfaces of a piezoelectric material. In a BAW device, three wave modes can propagate, namely, one longitudinal mode (embodying longitudinal waves, also called compressional/extensional waves), and two shear modes (embodying shear waves, also called transverse waves), with longitudinal and shear modes respectively identifying vibrations where particle motion is parallel to or perpendicular to the direction of wave propagation. The longitudinal mode is characterized by compression and elongation in the direction of the propagation, whereas the shear modes consist of motion perpendicular to the direction of propagation with no local change of volume. Longitudinal and shear modes propagate at different velocities. In practice, these modes are not necessarily pure modes as the particle vibration, or polarization, is neither purely parallel nor purely perpendicular to the propagation direction. The propagation characteristics of the respective modes depend on the material properties and propagation direction respective to the crystal axis orientations. The ability to create shear displacements is beneficial for operation of acoustic wave devices with fluids because shear waves do not impart significant energy into fluids.
Certain piezoelectric thin films are capable of exciting both longitudinal and shear mode resonance, such as hexagonal crystal structure piezoelectric materials including aluminum nitride (AlN) and zinc oxide (ZnO). To excite a wave including a shear mode using a standard sandwiched electrode configuration device, a polarization axis in a piezoelectric thin film must generally be non-perpendicular to (e.g., tilted relative to) the film plane. In biological sensing applications involving a liquid media, the shear component of a resonator is used because it is not damped completely by liquid loading. In this case, the piezoelectric material is grown with a c-axis orientation distribution that is non-perpendicular relative to a face of an underlying substrate to enable a bulk acoustic wave resonator structure to exhibit a dominant shear response upon application of an alternating current across electrodes thereof.
An electromechanical coupling coefficient is a numerical value that represents the efficiency of piezoelectric materials in converting electrical energy into acoustic energy for a given acoustic mode. Changing the c-axis angle of inclination for hexagonal crystal structure piezoelectric materials causes variation in shear and longitudinal coupling coefficients. FIG. 1 embodies a plot of shear coupling coefficient (Ks) and longitudinal coupling coefficient (Kl) each as a function of c-axis angle of inclination for AlN, although other piezoelectric materials show similar behavior. At certain angles (e.g., 46° and 90°) the longitudinal component is minimized and Kl has a zero value, and at certain angles (e.g., 0° and 67°) the shear component is minimized and Ks has a zero value. At all other angles of C-axis inclination, there exist both shear and longitudinal components of wave propagation. Devices built with C-axis angles that include both longitudinal and shear modes (e.g., at angles except for about 0°, 46°, 67°, and 90°) are referred to as quasi-shear mode devices. Quasi-shear mode acoustic resonator devices may be incorporated in fluidic devices providing sensing utility.
Under typical operating conditions, flows in microfluidic channels (also termed “microchannels”) are laminar. Fluids in laminar flow tend to follow parallel streamline paths, such that the chaotic fluctuations of velocity that tend to homogenize fluids in turbulent flows are absent. Multiple fluids introduced in a standard microchannel generally will not mix with each other, except at a common interface between the fluids via diffusion, and the diffusion process is typically slow compared with the flow of fluid along a principal axis of a microfluidic channel. The same principles that inhibit rapid mixing of fluids flowing under laminar conditions in a microfluidic channel also affect the distribution of analytes contained in one or more fluids flowing within a microfluidic channel. Fick's first law of diffusion states that flux moves from regions of high concentration to regions of low concentration. Secondarily, the flux rate is proportional to the concentration gradient difference. In a volume of fluid containing an analyte and advancing in a horizontal direction through a microfluidic channel having functionalization material arranged along a bottom surface of the channel, the fluid volume may be modeled as a moving “stack” of horizontal fluid layers. Even if it is assumed that analyte concentration is constant in each layer of the stack forming the fluid volume upon entering the microfluidic channel, following passage of the fluid volume over the functionalization material, a lowermost fluid layer of the stack will exhibit reduced or depleted analyte concentration due to binding of analyte with the functionalization material. But since diffusion is slow in a direction perpendicular to the direction of fluid flow through the microfluidic channel, and analyte needs to diffuse to a surface bearing functionalization material to bind, analyte present in fluid layers other than the lowermost fluid layer may not be available for binding with the functionalization material along the bottom surface of the channel within a reasonable period of time. Therefore, analyte concentration may remain stratified within the channel until diffusion occurs. Additionally, large analyte molecules may require a long time to bind with functionalization material.
Thus, conventional biochemical sensing devices may suffer from inconsistent distribution of target species in a sample and/or a low rate of analyte binding that may extend the time necessary to complete measurement of a particular sample. Accordingly, there is a need for fluidic devices incorporating BAW resonator structures, such as for biosensing or biochemical sensing applications, that overcome limitations associated with conventional devices.